Recent advances in nanotechnology and biology include the development of functional nanoparticles (electronic, optical, or magnetic) that are covalently linked to biological molecules such as peptides, proteins, and nucleic acids (See, e.g., Whaley et al., 2000; Bruchez et al., 1998; Chan et al., 1998; Mattoussi et al., 2000; Mitchell et al., 1999; Pathak et al., 2001; Elghanian et al., 1997; Reynolds et al., 2000; Mirkin et al., 1996; Storhoff et al., 1999; Alivisatos et al., 1996; Dubertet et al., 1996). Due to their size-dependent properties and dimensional similarities to biomolecules, these bioconjugates are well suited as contrast agents for in vivo magnetic resonance imaging (MRI)(Josephson et al., 1999; Bulte et al., 1997; Bulte et al., 2001), as luminescent dyes for cellular imaging (Han et al, 2001; Nicewarner-Pena et al., 2001), and as molecular beacons for ultrasensitive optical imaging and detection (Tyagi and Kramer, 1996; Tyagi et al., 1998; Sokol et al., 1998). However, the great potential of magnetic nanoparticles and their bioconjugates as efficient, high throughput and versatile tools for in vivo targeting, sensing and tracking of nucleic acids and proteins has just begun to be exploited (Weissleider, 2001; Shen et al., 1993; Hogemann et al., 2000; Weissleider and Mahmood, 2001).
The ability to monitor and quantify the level of gene expression in living cells in real time can provide important information concerning the production, temporal and spatial processing, localization, and transport of specific mRNA in different conditions. This new type of information could potentially revolutionize medical diagnostics and therapeutics. Technologies currently available for analysis and quantification of gene expression such as real-time RT-PCR, Northern blotting, expressed sequence tag (EST), serial analysis of gene expression (SAGE) and DNA microarrays are powerful tools for in vitro studies; however, they are not capable of quantifying gene expression in living cells. There is a clear need to develop molecular probes that can recognize target mRNA in living cells with high specificity and instantaneously convert such recognition into a measurable signal with a high signal-to-background ratio.
With recent advances in functional genomics and proteomics, there is an urgent need to develop technologies and strategies to detect proteins, multi-protein complexes and protein-protein interactions in vitro. For example, there is a lack of novel labeling reagents in performing in vivo protein expression and interaction studies. There exist few tools that can be used to visualize protein-protein interaction inside cells in deep tissue, and to track the interaction between drug molecules and the associated protein targets. There is virtually no existing method that enables studies of the dynamics of different intra-cellular molecular processes in deep tissue. This is a daunting challenge in that none of the many approaches under development to address these problems have yet demonstrated compelling promise—even as generally effective laboratory-scale small animal models.
Hair-pin oligonucleotide probes have enhanced specificity in detecting single-stranded RNA or DNA targets (Tsourkas et al, 2003). In particular, molecular beacons are a class of fluorescence-quenched hair-pin nucleic acid probes that can be used in a quantitative fashion; these probes fluoresce upon target recognition (i.e., hybridization) with potential signal enhancement of greater than 200 under ideal conditions. Structurally, they are dual-labeled oligonucleotides with a reporter fluorophore at one end and a dark quencher at the opposite end (Tyagi and Kramer; 1996). They are designed to have a target-specific probe sequence positioned centrally between two short self-complementary segments which, in the absence of target, anneal to form a stem-loop hairpin structure that brings the fluorophore in close proximity with the quencher. In this configuration the molecular beacon is in the “dark” state (Bernacchi and Mely, 2001). The hairpin opens upon hybridization with a complementary target, physically separating the fluorophore and quencher. In this configuration the molecular beacon is in the “bright” state. Transition between dark and bright states allows for differentiation between bound and unbound probes and transduces target recognition into a fluorescence signal (Matsuo, 1998; Liu et al., 2002).
The unique target recognition and signal transduction capabilities of molecular beacons have led to their application in many biochemical and biological assays including quantitative PCR (Vogelstein and Kinzler, 1999; Chen and Mulchandani, 2000), protein-DNA interactions (Fang et al., 2000; Li et al., 2000), multiplex genetic analysis (Marras et al., 1999; de Baar et al., 2001), and the detection of mRNA in living cells (Matsuo 1988; Sokol et al., 1998; Molenaar 2001). However, false-positive signals due to protein-beacon interaction and nuclease-induced beacon degradation significantly limit the sensitivity of the in vivo applications (Mitchell, 2001). The thermodynamic and kinetic properties of molecular beacons are dependent on its structure and sequence in complex ways (Bonnet et al. 1999; Kuhn et al., 2002). Moreover, the signal-to-background ratio in target detection is dependent not only on design (length and sequence of the stem and probe) but also on the quality of oligonucleotide synthesis and purification (Goddard et al., 2000; Bonnet et al., 1998) and the assay conditions employed. Although molecular beacons, with the dual FRET (fluorescence resonance energy transfer) design (Tsourkas et al, 2003a) and backbone modifications (Tsourkas at el, 2002b), can be used for sensitive and quantitative detection of mRNA in living cells, they are optical-based probes and therefore cannot be used for deep tissue imaging of RNA expression due to light absorption and scattering.
Compared with micrometer sized magnetic particles and chelates of paramagnetic ions such as gadolinium diethylenetriaminopentaacetic acid (Gd-DTPA), magnetic nanoparticles are much more efficient as relaxation promoters and their effect on the relaxivities of water is measurable even at nanomolar concentrations, as demonstrated theoretically (Koenig and Kellar, 1995) and experimentally (Bulte et al, 1992; Le Duc et al, 2001; Josephson, Perez and Weissleder, 2001; Perez et al, 2002). Specifically, the use of ferromagnetic and superparamagnetic nanoparticles as contrast agents can induce a more than 10 fold increase in proton relaxivities (Coroiu et al, 1999). Based on the enzyme-cleavable contrast agents reported by Meade and coworkers (Huber et al, 1998; Louie et al, 2000), magnetic relaxation by gadolinium atoms is a short-range effect. The presence of a carbohydrate cap (less than 1 nm thick) prevents water molecules from contacting the contrast agent Gd (HP-DO3A). In contrast, superparamagnetic nanoparticles (2-5 nm core size) are often coated by a thick dextran layer (10-20 nm) (Josephson, Perez and Weissleder, 2001), and the encapsulated nanoparticles are highly efficient contrast agents, yielding considerably stronger effects than gadolinium compounds.
Superparamagnetic iron oxide nanoparticles have been extensively studied for different applications including high-density magnetic storage, magnetic nanostructures, catalytic and separation processes (Dyal et al., 2003), contrast agents for magnetic resonance imaging applications (Zhao et al., 2002), functional cell labeling for in vivo tracking of stem cells and tumor progression (Lewin et al., 2000; Bulte et al., 2001), cell and DNA sorting applications (Dressman et al., 2003), drug delivery applications (Lanza et al.), and cellular mechanics studies. Most of these applications require water-soluble, stable, monodispersed, and uniformly sized magnetic nanoparticles. In addition, there is growing interest in applications of magnetic nanoparticles for intracellular molecular imaging. What is needed in the art is to have biocompatible magnetic nanoparticles with a size comparable to the molecular targets (approximately 2-15 nm), with multifunctional ligands. These ligands need to facilitate efficient delivery of nanoparticles, enable specific cell-type recognition, provide suitable biological molecules (antibodies, oligonucleotides etc.) for intracellular targeting, and generate specific MRI contrast via clustering of magnetic nanoparticles or other molecular switch mechanisms.
A variety of coating strategies done both in situ and post-synthesis have been developed for stabilization of magnetic nanoparticles. Post-synthesis coating processes include monolayer ligands (Liu and Xu, 1995; Yee et al., 1999), polymer coatings (Burke et al., 2002; Harris et al., 2003), and silica coatings (Butterworth et al., 2001; Santra et al., 2001; Lu et al., 2002). The monolayer ligand coatings rely on interactions of chemical groups from both the ligand and nanoparticle for effective adsorption or chemisorption of the ligand on the surface. These coatings tend to have limited colloidal stability due to weak steric hindrance in preventing aggregation and have limited opportunity for functionalization. Residual surfactant on the surface of nanoparticles can result in inefficient or incomplete coatings with this approach. The amphiphilic polymer-based or silica-based coating approach often results in multilayer coatings (Lu et al., 2002), making the coating process difficult to control and resulting in a heterogeneous sample, and sometimes in the encapsulation of dimers or trimers of nanoparticles in the same shell (Lu et al., 2002). To improve ease of functionalization and colloidal stability, silica-based coatings are often supplemented with polymers such as polyethylene glycol (Butterworth et al., 2001). In addition to post-synthesis approaches, some in situ coating approaches have been developed, such as the commonly used dextran coatings for MIONs (Kim et al., 2003). These coating processes also lead to multilayer coatings, which can result in a heterogeneous sample. In situ synthesis conditions limit the amount of functionalization that can be achieved.
Recent research has identified several small regions (10-16 amino acids) of proteins called protein transduction domains (PTDs) that possess the ability to traverse biological membranes efficiently in a process termed protein transduction (Lewin et al., 2000; Wadia et al., 2002; Becker-Hapak et al., 2001). Examples of PTDs are HIV-1 Tat, HSV VP22, and ANTP. It was shown that transduction occurs in a receptor- and transporter-independent fashion that appears to target the lipid bilayer directly. Thus, in principle and practice, all cell types appear transducible. Proteins and compounds that are covalently linked to PTDs have proven to be useful in answering specific biological questions where other methods fail. Moreover, PTD fusion proteins have now been introduced into mice and exhibit delivery of active enzyme to all tissues, including across the blood-brain barrier. Recent research shows that magnetic particles as large as 40 nm and liposomes as large as 100 nm can be delivered into cells and tissues.
Traditionally, MRI subjects a tissue to a uniform magnetic field to analyze whether certain features are present or atypical in the subject. MRI uses changes in the angular momentum or “spin” of atomic nuclei of certain elements to show locations of those elements within matter. In particular, a static magnetic field (typically between 0.2 and 4 Tesla) acts on the nuclei of atoms with fractional spin quantum numbers and polarizes them into alignment within the magnetic field. During MRI measurements, radio-frequency pulses of a given resonance energy are applied that flip the nuclear spins and disturb the orientation distribution. The radio-frequency (RF) pulses are in the form of a second, oscillating, RF magnetic field having a particular frequency referred to in the art as a resonant or Larmor frequency. This frequency is equal to the rate that the spins rotate or precess. When the magnetic field is turned off, the nuclei return (relax) to the initial state in a time dependent exponential fashion, thus giving signals which are electronically processed into recordable data. When the signals are spatially differentiated and of sufficient level, the data can be organized and displayed, allowing direct visualization of tissues and organs of a patient. Contrast agents, such as free metal ions, chelates, or insoluble metal compounds have been described for use in enhancing intrinsic contrast in MR imaging. Such metals may include, for example, iron, cobalt, zinc, cadmium, chromium, copper, nickel, manganese, and gadolinium. Magnetic nanoparticles induce a measurable change in the relaxation times T1 and T2 (and T2*) of surrounding water molecules that can be visualized using MRI. In addition, the coupling or clustering of magnetic nanoparticles generates a different effect on water. These effects of individual or clustered magnetic nanoparticles on water can be differentiated using MRI. However, traditional MRI contrast agents, as well as the use of magnetic nanoparticles with MRI to target cell surface receptors, are not suitable for detecting differential levels of gene expression inside living cells.
Therefore, there is a strong need in the art to provide multifunctional magnetic nanoparticle probes that are suitable for in vivo delivery and that have both signal transduction and target recognition properties. There is also a need for such compositions and methods that can be used for detection of genetic transcription in vivo and in specific cell types Furthermore, there is a need for such compositions and methods that can be used for both optical intra-cellular imaging and deep-tissue molecular imaging using MRI. Moreover, there is a need for such compositions and methods for the rapid and ultrasensitive in vivo detection of diseased cells; virus infection; the rapid detection of internal injury such as traumatic brain injury; and molecular imaging and monitoring of drug activities and therapeutic effects.